A normal ear transmits sounds as shown in FIG. 1 through the outer ear 101 to the tympanic membrane (eardrum) 102, which moves the bones of the middle ear 103 (malleus, incus, and stapes) that vibrate the oval window and round window openings of the cochlea 104. The cochlea 104 is a long narrow duct wound spirally about its axis for approximately two and a half turns. It includes an upper channel known as the scala vestibuli and a lower channel known as the scala tympani, which are connected by the cochlear duct. The cochlea 104 forms an upright spiraling cone with a center called the modiolar where the spiral ganglion cells of the acoustic nerve 113 reside. In response to received sounds transmitted by the middle ear 103, the fluid-filled cochlea 104 functions as a transducer to generate electric pulses which are transmitted to the cochlear nerve 113, and ultimately to the brain.
Hearing is impaired when there are problems in the ability to transduce external sounds into meaningful action potentials along the neural substrate of the cochlea 104. To improve impaired hearing, auditory prostheses have been developed. For example, when the impairment is related to operation of the middle ear 103, a conventional hearing aid may be used to provide acoustic-mechanical stimulation to the auditory system in the form of amplified sound. Or when the impairment is associated with the cochlea 104, a cochlear implant with an implanted electrode contact can electrically stimulate auditory nerve tissue with small currents delivered by multiple electrode contacts distributed along the electrode.
FIG. 1 also shows some components of a typical cochlear implant system which includes an external microphone that provides an audio signal input to an external signal processor 111 where various signal processing schemes can be implemented. The processed signal is then converted into a digital data format, such as a sequence of data frames, for transmission into the implant 108. Besides receiving the processed audio information, the implant 108 also performs additional signal processing such as error correction, pulse formation, etc., and produces a stimulation pattern (based on the extracted audio information) that is sent through an electrode lead 109 to an implanted electrode array 110. Typically, this electrode array 110 includes multiple stimulation contacts 112 on its surface that provide selective stimulation of the cochlea 104.
The electrode array 110 contains multiple electrode wires embedded in a soft silicone body referred to as the electrode carrier. The electrode array 110 needs to be mechanically robust, and yet flexible and of small size to be inserted into the cochlea 104. The material of the electrode array 110 needs to be soft and flexible in order to minimize trauma to neural structures of the cochlea 104. But an electrode array 110 that is too floppy tends to buckle too easily so that the electrode array 110 cannot be inserted into the cochlea 104 up to the desired insertion depth.
Typically, the electrode wires within the electrode array 110 have a homogenous overall shape from one end to the other: either generally straight, repeating coiled loops, or recurring wave shapes. As shown in FIG. 17, the bend radius of the electrode array 110 becomes ever smaller as it is inserted more deeply into the cochlea. So the electrode array 110 should have non-uniform and non-homogeneous mechanical properties (e.g., bending and flexing) to accommodate the complex path that it must take, and also for maintaining biological compatibility with the surrounding tissue of the cochlea 104.
In addition, present cochlear implant (CI) systems possess numerous stimulation contacts 112 along the electrode array 110 for achieving a frequency distribution and resolution that mimics natural human hearing as far as possible. As the technology advances it is likely that an increasing number of frequency bands will need to be supported by the CI systems for providing an even finer pitched hearing. Consequently, more and more wires and stimulation contacts 112 will have to be placed within the electrode array 110, whose dimensions are restricted by the very limited space in the cochlea 104. In general, it can be said that the more channels (i.e. wires and contacts) an electrode array 110 contains, the more rigid it will be due to the higher amount of metal structures within it.
A trade-off needs to be made between a certain stiffness of the electrode array 110 which allows insertion into the cochlea 104 up to the desired insertion depth without the array buckling, and certain flexibility of the electrode array 110 which keeps mechanical forces on the lateral wall of the scala tympani of the cochlea 104 low enough.
Recent developments in CI electrode array designs and surgical techniques are moving towards minimal trauma implantations. For preservation of residual hearing it is of particular importance to preserve the natural intra-cochlear structures. Therefore, the size and mechanical characteristics of the electrode array are critical parameters for the best patient benefit. Some electrode array designs are pre-curved, though a drawback of that approach is that a special electrode insertion tool is needed which keeps the electrode array straight until the point of insertion.
As documented by Erixon et al., Variational Anatomy of the Human Cochlea: Implications for Cochlear Implantation, Otology & Neurotology, 2008 (incorporated herein by reference), the size, shape, and curvature of the cochlea varies greatly between individuals, meaning that a CI electrode array must match a wide range of scala tympani (ST) geometries. Furthermore, recently published research by Verbist et al., Anatomic Considerations of Cochlear Morphology and Its Implications for Insertion Trauma in Cochlear Implant Surgery, Otology, & Neurotology, 2009 (incorporated herein by reference) has shown that the human ST does not incline towards the helicotrema at a constant rate, but rather there are several sections along the ST where the slope changes, sometimes even becoming negative (i.e. downwards). The location and grade of these changes in inclination were also found to be different from individual to individual. Consequently, CI electrode arrays should be highly flexible in all directions in order to adapt to individual variations in curvature and changes in inclination of the ST for minimal trauma implantation.
Present day CI electrode arrays require considerable amount of hand assembly during manufacturing. Single thin platinum wires covered with a thin electrical insulation must be cut to size and manipulated without compromising the insulation. The wires must be stripped of insulation at the ends and welded to small thin platinum foils that act as stimulation contacts. Each wire must be individually placed inside a mold and assembled in a multi-channel structure before being silicone injection molded. Demolding of long electrodes must take place without causing damage to the structure.
Some rejects inevitably occur during manufacturing due to open or short circuits between wires, or poor welding to the contacts. Silicone overflow on contact surfaces may cause further rejects. The process of making electrodes is extremely labor intensive and a considerable percentage of rejected electrodes is unavoidable since maintenance of acceptable quality is difficult. In addition, the manual work is very operator dependent and difficult to specify in adequate detail to give reproducible results. Hand-made devices may therefore unintentionally and undesirably be subject to significant variations in performance. Furthermore, manual work is linked with extensive and time-consuming training of personnel and manual production may in general not be financially competitive.
It would therefore be desirable to have a streamlined method for making implant electrodes using an automated process. The requirements as to number of stimulation channels, size, and mechanical properties constitute a challenging problem for traditional and modern electrode manufacturing techniques. U.S. Pat. No. 6,374,143 by Berrang et al. (“Berrang”, incorporated herein by reference) presents a process for fabricating thin-film CI electrodes by encapsulating platinum structures between two polymer films. This process can be automated and thus attempts to address the problem of a lacking streamlined electrode manufacturing as described above. In the same patent, folding is suggested for miniaturization of an electrode array in order to pack the many metal wires into a smallest possible space. U.S. Pat. No. 7,085,605 by Bluger et al. (“Bluger”, incorporated herein by reference) discloses a similar method for an implantable medical assembly. WO2008/011721 by Spruit (“Spruit”, incorporated herein by reference) proposes stacking of several individual assembly layers for essentially achieving the same compact structure. Other methods for manufacturing a thin-film CI electrode include ink-jet printing of platinum ink onto a polymer film, as suggested by U.S. patent application Ser. No. 12/787,866, filed May 26, 2010 (incorporated herein by reference).
As the number of stimulation channels increases, an increasing number of folded or stacked layers is needed for electrically insulating the conducting metal wires from each other. One basic mechanical property of the described (folded or stacked) assemblies is the highly inhomogeneous bending characteristics in different directions, mainly caused by the geometry of the assembly layers containing the wires. The cross-section of these layers is rectangular in shape and therefore has a preferred bending direction. Existing and suggested CI electrode arrays based on the thin-film technology were therefore designed to be highly bendable in the direction of the ST curvature around the modiolus, but far less flexible in the plane parallel to the modiolus. As explained earlier, these characteristics are generally not desirable in CIs since they should be highly bendable in all directions to lower the risk of implantation trauma.
U.S. Pat. No. 5,964,702 (“Grill”, incorporated herein by reference) describes stimulating peripheral nerves using cuff electrodes wound in a helical shape where the stimulation contact surfaces are opened inwards towards the internal lumen of the helical shape. WO93/20887 (“Grill WO”, incorporated herein by reference) describes a similar arrangement for thin film implant electrodes. Both Grill methods use a first layer of elastomer that is cured and stretched and then covered by second layer of elastomer so that the different mechanical tensions in the two elastomer layers cause the layered structure to curl into a helix. But in pacemaker electrodes, the size constraints, the number of electrically active channels, and the requirements to flexibility (for preservation of delicate tissues) are fundamentally different than for many specific implant applications such as CI electrodes. It is therefore a challenge to produce CI electrodes that make use of the highly flexible helical shaped wires.
U.S. Patent Publication 2010/0305676 (“Dadd,” incorporated herein by reference) describes winding the electrode wires in the extra-cochlear segment of the electrode lead in a helical shape to make that portion of the electrode lead stronger. Dadd is quite clear that such a helical portion does not extend into the intra-cochlear electrode array which needs to be much more flexible than the extra-cochlear lead in order to minimize trauma to the cochlear tissues when the array is inserted.
U.S. Patent Publication 2010/0204768 (“Jolly,” incorporated herein by reference) describes winding the individual electrode wires in the intra-cochlear electrode array in an elongated helical shape where each wire is separate and independent.